This invention relates to a method for determining the stroke volume and, hence, the cardiac output of a patient, as well as to a system that implements the method.
Accurate measurement of the cardiac output (CO) of a patient has proven to be a valuable diagnostic tool. Accordingly, several methods for determining CO have been developed, of which thermodilution, the direct oxygen Fick method, and the pulse contour method (PCM) are at present the most prevalent. These known methods for measuring CO, however, are affected by several drawbacks that greatly limit their application in the clinical setting as well as for purposes of research.
Measurement of CO using thermodilution, which is described, for example in Ganz, W. and Swan, H. J. C. (1972), xe2x80x9cMeasurement of blood flow by thermodilution,xe2x80x9d Am. J. Cardiol. 29, pp. 241-246, has become routine in the hemodynamic evaluation and management of critically ill patients. As is well known, this method is based on the law of conservation of energy and on the application of the Stewart-Hamilton equation, for which a number of conditions must be fulfilled. These conditions include complete mixing of the thermal indicator with blood, no loss of indicator within the dilution volume, and constant blood flow during the dilution time.
Inaccuracy in the determination of CO may result from the inconsistency of these assumptions in many clinical conditions. In particular, variability of blood flow may occur as a consequence of hemodynamic instability related to changes in heart rate, cardiac arrhythmia, valvular or congenital heart disease, and application of mechanical ventilation. Additional limitations of the thermodilution method are its invasiveness and the impossibility of monitoring CO beat-to-beat in critical conditions and during the course of acute pharmacological interventions.
The direct oxygen Fick approach is the standard reference technique for CO measurement. See, for example, Fagard, R. and Conway, 3 (1990), xe2x80x9cMeasurement of cardiac output: Fick principle using catheterization,xe2x80x9d Eur. Heart J. 11, Suppl. I, pp. 1-5. According to the Fick principle, CO can be determined by the ratio of oxygen uptake to the difference in oxygen content between arterial and mixed venous blood. The validity of the principle depends upon the assumption that pulmonary blood flow is approximately identical to systemic blood flow and that the lungs themselves do not extract oxygen. Although this method appears to be the most accurate among those currently available, its use is limited by a series of practical problems. These problems include the need for right heart catheterization to obtain truly mixed venous blood, the assumption of the availability of appropriate analytical techniques for measuring oxygen uptake and content, and the attainment of a steady state in which apparent oxygen consumption matches tissue oxygen utilization. The fulfillment of these conditions makes the method unsuitable for repeated measurements and, consequently, not apt to follow rapid changes in flow over time.
The pulse contour method (PCM), which has been developed from an original idea by J. A. Herd et al. dating back to 1864 and from a theory commonly referred to as the xe2x80x9cWindkesselxe2x80x9d (German for xe2x80x9cair chamberxe2x80x9d) theory of Franck (Franck O., 1930), derives CO from the arterial pressure pulse wave. The PCM method is based on the existence of a relationship between the volume of blood expelled by the left ventricle (LSV) or the volume of blood expelled by the right ventricle (RSV) and the area under the pressure curve P(t). Unlike the thermodilution and Fick methods, which measure mean CO over a limited time span, the PCM operates on a beat-to-beat basis. The primary assumption of PCM is that the pressure rise during systole is related, in a complex way, to the systolic filling of the aorta and proximal large arteries. Various approaches have therefore been devised to approximate, by means of different models of the arterial system, the relationship between aortic pressure and flow.
One of the most famous models used in PCM was developed by Wesseling and his co-workers and is described in, among many other references:
Wesseling, K. H., Dc Wit, B., Weber, J. A. P. and Smith, N. T. (1983), xe2x80x9cA simple device for the continuous measurement of cardiac output. Its model basis and experimental verification,xe2x80x9d Adv. Cardiol. Phys. 5, Suppl II, pp.16-52;
Wesseling, K. H., Jansen, J. R. C., Settels, J. J. and Schreuder, J. J. (1993), xe2x80x9cComputation of aortic flow from pressure in humans using a nonlinear, three-element model,xe2x80x9d J. Appl. Physiol. 74, pp. 2566-2573;
Jansen, J. R. C., Wesseling, K. H., Settels, J. J. and Schreuder, J. J. (1990), xe2x80x9cContinuous cardiac output monitoring by pulse contour during cardiac surgery,xe2x80x9d Eur. Heart J. 11, Suppl 1, pp. 26-32;
Sprangers, R. L., Wesseling, K. H., lmholz, A. L., Imholz, B. P. and Wieling, W. (1991), xe2x80x9cInitial blood pressure fall on stand up and exercise explained by changes in total peripheral resistance,xe2x80x9d J. Appl. Physiol. 70, pp. 523-530;
Jellema, W. T., Imholz, B. P. M., van Goudoever, J., Wesseling, K. H. and van Lieshout, J. J. (1996), xe2x80x9cFinger arterial versus intrabrachial pressure and continuous cardiac output during head-up tilt testing in healthy subjects,xe2x80x9d Clin. Sci. 91, pp.193-200;
Stock, W. J., Baisch, F., Hillebrecht, A., Schulz, H. and Karemaker, J. M. (1993), xe2x80x9cNoninvasive cardiac output measurement by arterial pulse analysis compared to inert gas rebreathing,xe2x80x9d J. Appl. Physiol. 74, pp. 2687-2693;
Harms, M. P. M., Wesseling, K. H., Pott, F., et al. (1999), xe2x80x9cContinuous stroke volume monitoring by modelling flow from non-invasive measurement of arterial pressure in humans under orthostatic stress,xe2x80x9d Clin. Sci. 97, pp. 291-301;
Houtman, S., Oeseburg, B. and Hopman, M. T. E. (1999), xe2x80x9cNon-invasive cardiac output assessment during moderate exercise: pulse contour compared with C02 rebreathing,xe2x80x9d Clin. Physiol. 19, pp. 230-237;
Jellema, W. T., Wesseling, K. H., Groeneveld, A. B. J, Stoutenbeek, C. P., Thjis, L. G. and van Lieshout, J. J. (1999), xe2x80x9cContinuous cardiac output in septic shock by simulating a model of the aortic input impedance. A comparison with bolus injection thermodilution,xe2x80x9d Anesthesiology 90, pp.1317-1328;
Langewouters, G. J., Wesseling, K. H. and Goedhard, W. J. A. (1984), xe2x80x9cThe static elastic properties of 43 human thoracic and 20 abdominal aortas in vitro and the parameters of a new model,xe2x80x9d J. Biomech. 17, pp. 425-435; and
Stock, W. J., Stringer, R. C. 0. and Karemaker, J. M. (1999), xe2x80x9cNoninvasive cardiac output measurement in orthostasis: pulse contour analysis compared with acetylene rebreathing,xe2x80x9d J. Appl. Physiol. 87, pp. 2266-2273.
The Wesseling method is based on a model of the elastic properties of the aorta and has been found to be satisfactory under certain hemodynamic circumstances. According to the xe2x80x9cModelflowxe2x80x9d method developed by Wesseling and coworkers, uncalibrated values of CO are obtained by relating the area under the pulsatile systolic portion of the pressure wave to parameters derived from a nonlinear three-element model of the arterial system. In PCM, in order to establish a relationship between pressure and flow, the mechanical properties of the arteries, as function of arterial pressure, are approximated either by several empirical formulae or by using a model based on age- and sex-predicted values not directly pertaining to the subject under study.
The three elements of the model used in Modelflow are aortic characteristic impedance (i.e., the relationship between the rise of pressure in the aortic root in opposition to the flow of blood ejected from the left ventricle), arterial compliance (i.e., the relationship between changes in blood volume and changes in pressure in the aorta), and peripheral vascular resistance (i.e., the relationship between mean pressure and mean flow). The first two elements of the modelxe2x80x94impedance and compliancexe2x80x94depend mostly on the elastic properties of the aorta. In Modelflow, these elements are predicted by an experimentally derived arctangent function that relates aortic pressure and cross-sectional area; however, this prediction has, as required input variables, the age and sex of the subject. The third elementxe2x80x94vascular resistancexe2x80x94is derived from the model simulation and is calculated and updated for the next heartbeat by the ratio between mean pressure and the computed flow.
Compared to other methods, the major advantage of PCM is the ability to monitor CO beat-to-beat over prolonged periods without the need of an indwelling pulmonary artery catheter for recording temperature changes or for blood sampling. Indeed, measurements can be derived from the pressure recordings in a systemic peripheral artery or even from the pressure signal detected noninvasively at the finger. The results of the method rely on the aortic pressure-cross-sectional area relationship, which is approximated from unrelated in vitro measurements on segments of human thoracic aorta. To obtain absolute values of CO, it is then necessary to determine, at least once for each patient, a calibrating factor of the model parameters by comparison of the PCM result with an absolute CO estimatexe2x80x94without such calibration, PCM can provide only relative changes in CO. The need for comparison with a reference method greatly limits the usefulness of PCM since the calibrating technique is either invasive (e.g., thermodilution) or cumbersome (e.g., inert gas rebreathing) and it must be repeatedly applied when changes in the experimental procedure, which may alter the physical properties of the arteries, are induced.
The estimates of CO by Modelflow thus depend more on fixed predicted parameters than on actual measurements obtained from the subject under evaluation. In fact, the parameters to measure CO derived directly from the pressure wave are limited to pulsatile systolic area, mean blood pressure, and heart rate. Other parameters that characterize the elastic properties of the arteries and that can be derived from the shape of the pressure curve, such as the time of attainment of peak systolic pressure, the presence of sudden slope changes, and the length of the diastolic phase, are not taken into consideration for the computation of CO. As a consequence, different forms of the pressure signal and different end-diastolic pressure levels can result in pulsatile systolic areas with comparable integral value. It is likely, however, that pressure waves with markedly different contours and end-diastolic pressure levels may reflect definite differences of arterial vessels physical characteristics even though they have similar pulsatile systolic areas. As such, accurate computation of CO using Modelflow and related methods is highly dependent on the measurement of a calibrating factor derived by comparison with an independent standard reference method, rather than on the actual pressure wave morphology.
What is needed is therefore a method for measuring accurately an absolute value of CO, that is, which does not require calibration using some other absolute method, and that can do so continuously for prolonged periods. The method should preferably be able to accomplish this without the level of invasiveness of, for example, thermodilution methods, and it should not depend on such patient-specific parameters such as weight, height, body surface area, sex, age, the diameter of the aorta of the patient, etc. This invention provides such a method, and a related system for implementing it.
The invention provides a method for measuring cardiac output (CO) of a patient, as well as a system that implements the method. According to the invention, arterial blood pressure is sensed and is converted to a pressure signal. An estimate of stroke volume is then calculated as a function only of selected characteristics of the sensed pressure signal and of predetermined, patient unspecific constants. The invention then calculates an estimate of CO as a function of the estimated stroke volume and a current heart rate value.
In the preferred embodiment of the invention, the estimate of stroke volume is calculated by calculating an area (A) under the entire pressure signal, including both pulsatile and non-pulsatile portions of the pressure signal, over a cardiac cycle; estimating selected impedance values from the pressure signal; and calculating the estimate of stroke volume as a function of the ratio between the calculated area and the estimated selected impedance values.
As part of calculating the estimate of stroke volume, the invention preferably also calculates a mean pressure value of the pressure signal and then corrects the estimated stroke volume as a predetermined function of the mean pressure value and of a reference pressure.
The preferred embodiment of the invention detects the times and corresponding pressure values of a systolic peak and of a dicrotic notch in the pressure signal. A second derivative of the pressure signal is then evaluated between the systolic peak and the dicrotic notch. The time and corresponding pressure value of at least one intermediate point in the pressure signal are then detected between the systolic peak and the dicrotic notch at which the second derivative has an extreme value and at least one of the selected impedance values is then estimated as a predetermined function of the time and corresponding pressure value of the intermediate point.
A systolic peak pressure Psys, a diastolic pressure Pdia, and a dicrotic pressure Pdic are preferably detected and the estimated stoke volume is then preferably scaled by a factor proportional to the ratio between the difference between Pdia and Pdic and the difference between Psys and Pdia.
In addition to detecting a dicrotic notch in the pressure signal, in some embodiments of the invention, a post-dicrotic first derivative of a post-dicrotic portion of the pressure signal is also evaluated at times after the dicrotic notch. In these cases, the time and corresponding pressure value of at least one local maximum pressure are detected in the post-dicrotic portion of the pressure signal and at least one of the selected impedance values is estimated as a predetermined function of the time and corresponding pressure value of the local maximum pressure.
Using the preferred embodiment of the invention, the CO estimate may be calculated based on the pressure signal during a single cardiac cycle.
In the preferred embodiment of the invention, the pressure signal is uncalibrated, whereby the steps of calculating the estimate of the stroke volume and of calculating the estimate of CO are independent of external calibration.
The invention thus provides for direct calculation of cardiac flow (or, equivalently, cardiac output) and arterial impedance (Z) from a measured pressure signal. The pressure signal may be measured either invasively, for example, in the ascending aorta, or in the pulmonary, femoral, brachial, or radial artery, or non-invasively, for example from the arteriole of the finger using a cuff meter. A composite impedance value Ztot of the pressure signal is calculated directly on the basis of characteristics of various resonance points, preferably by means of an analysis of the first and second time derivatives of the pressure signal.
In determining stroke volume SV, from which the invention calculates cardiac output, the invention takes into account both pulsatile and non-pulsatile (continuous) components of the recorded pressure signal. Accordingly, in calculations of SV, the invention considers the area A under the entire pressure signal, that is, the pulsatile portion above the diastolic pressure as well as the continuous portion below the diastolic pressure. Moreover, with respect to the composite impedance Ztot, in addition to the pulsatile portion between the time of diastoly and the dicrotic notch, the invention also takes into consideration the influence of the non-pulsatile, continuous portion of the pressure curve that occurs after the dicrotic notch.
The invention is therefore able to calculate the cardiac flow with no need to calibrate the recorded pressure signal, and no need to incorporate patient-specific, anthropometric data. Rather, the invention determines SV exclusively from an analysis of the characteristics of the pressure wave itself. These characteristics include not only the xe2x80x9cprinciplexe2x80x9d balancing points (systolic and dicrotic points) of pressure of ventricularly ejected blood, but also of additional points of balance.